Sensing platform for transduction of information

ABSTRACT

Aspects of a biosensor platform system and method are described. In one embodiment, the biosensor platform system includes a fluidic system and tunneling biosensor interface coupled to the fluidic system. The tunneling biosensor interface may include a transducing electrode array having at least one dielectric thin film deposited on an electrode array. The biosensor platform system may further include processing logic operatively coupled to the transducing electrode array. In operation, the application of an electromagnetic field at an interface between an electrode and an electrolyte in the system, for example, may result in the transfer of charge across the interface. The transfer of charge is, in turn, characterized by electromagnetic field-mediated tunneling of electrons that may be assisted by exchange of energy with thermal vibrations at the interface. Various analytes, for example, and other compositions can be identified by analysis of the transfer of charge.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.14/455,205, entitled “SENSING PLATFORM FOR TRANSDUCTION OF INFORMATION,”filed Aug. 8, 2014, which claims the benefit of U.S. ProvisionalApplication No. 61/864,072, filed Aug. 9, 2013, the entire contents ofboth of which applications are hereby incorporated herein by reference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with government support under contractN66001-11-1-4111 awarded by the Defense Advanced Research ProjectsAgency. The government has certain rights in this invention.

BACKGROUND

In a variety of applications, the detection and identification ofcertain chemical or molecular species is desired. For example, it may bedesirable to identify small molecule analytes, such as amino acids andmetallic ions, as well as relatively large proteins, such as DNA andRNA. In particular, the detection of biomarkers in biological samples isimportant for disease detection, disease analysis, and disease pathwayinvestigation. Further, the detection of contaminants in environmentalsamples, such as in water, is important for homeland security, publicsafety, and environmental welfare.

BRIEF DESCRIPTION OF THE DRAWINGS

For a more complete understanding of the embodiments described hereinand the advantages thereof, reference is now made to the followingdescription, in conjunction with the accompanying figures brieflydescribed as follows:

FIG. 1A illustrates an example schematic diagram of an electrochemicalinterface with characteristic length scales to determine the nature of acharge transfer reaction according to aspects of the embodiments.

FIG. 1B illustrates an example regime, as an interface charge density,within which transduction of molecular vibration modes is possibleaccording to aspects of the embodiments.

FIG. 2A illustrates an example of weak coupling between electronicenergy and nuclear-vibrational states.

FIG. 2B illustrates an example of strong coupling between electronicenergy and nuclear-vibrational states.

FIG. 2C illustrates an example reaction free energy schematic for anadiabatic reaction case when an electron source and donor (initial andfinal electronic energy states) are strongly coupled according toaspects of the embodiments.

FIG. 2D illustrates an example reaction free energy schematic for anon-adiabatic reaction case when initial and final electronic energystates are weakly coupled according to aspects of the embodiments.

FIG. 3 illustrates an example schematic depicting the measurement of aflux of electrons crossing an electrified dielectric monolayer modifiedelectrochemical interface and the interaction of the tunneling electronswith an analyte co-located at the interface according to aspects of theembodiments.

FIG. 4 illustrates example design considerations for a transducinginterface and factors in view of quantum-mechanical to classicaltransition behavior according to aspects of the embodiments.

FIG. 5 illustrates example experimental data acquired from a sensinginterface representative of the sensitivity of the biosensor describedherein to a single atom isotope substitution.

FIGS. 6A-C illustrate example embodiments and geometries ofdielectric-film-modified nanoscale electrode-electrolyte interfaces andof nano-engineered interfaces according to aspects of the embodiments.

FIG. 7 illustrates an example biosensor platform according to certainaspects of the embodiments described herein.

FIG. 8 illustrates an enlarged view of an example sensor with electrodesarranged in an array according to aspects of the embodiments.

FIG. 9 illustrates processes of spectral data collection, referencedatabase collection, and analysis according to aspects of theembodiments.

FIG. 10 illustrates example tunneling barriers at metal-dielectric anddielectric-electrolyte interfaces according to aspects of theembodiments.

FIG. 11 illustrates sequential layering of high and low k-dielectricmaterials for a high-k dielectric insulator according to aspects of theembodiments.

FIG. 12 illustrates a magnetic tunneling film architecture that usesdifferentially oriented film magnetic moments to further restrictelectronic transition according to aspects of the embodiments.

FIGS. 13A-C illustrate example embodiments of gate-electrode systemsaccording to aspects of the embodiments.

FIG. 14 illustrates an example of three-electrode feedback suppressionof thermal noise for electronic transition measurements according toaspects of the embodiments.

The drawings illustrate only example embodiments and are therefore notto be considered limiting of the scope described herein, as otherequally effective embodiments are within the scope and spirit of thisdisclosure. The elements and features shown in the drawings are notnecessarily drawn to scale, emphasis instead being placed upon clearlyillustrating the principles of the embodiments. Additionally, certaindimensions may be exaggerated to help visually convey certainprinciples. In the drawings, similar reference numerals between figuresdesignate like or corresponding, but not necessarily the same, elements.

DETAILED DESCRIPTION OF THE EMBODIMENTS

As described above, the detection and identification of certain chemicalor molecular species is desired in various fields and applications. Inparticular, the detection of biomarkers in biological samples isimportant for disease detection, disease analysis, and disease pathwayinvestigation. Further, the detection of contaminants in environmentalsamples, such as in water, is important for homeland security, publicsafety, and environmental welfare.

Using conventional means and methods, certain chemical and molecularspecies may be identified. The identification of these species may beachieved using bioassays, electronic systems, or combinations thereof,for example. Typically, a bioassay may indirectly detect analytes bymeasuring various molecular interactions. Some bioassays detect analytesby activating a label that is covalently attached to a binding partnerupon analyte binding to a bait molecule. Other bioassays measure analytebinding of an immobilized bait molecule to a solid substrate and changesin charge, refractive index, or mass change at an interface between thesolid substrate and liquid sample. In various forms, electronic systemsmay rely upon alterations in current, voltage, or charge to indirectlydetect, qualify, and quantify chemical analytes. It should beappreciated, however, that the demand for a low-cost and field-usefriendly means or method to identify and detect low concentrationanalytes has resulted in ongoing efforts to improve the functionalityand practicality of chemical and molecular detecting devices.

A good platform for detecting biological threats should be able toidentify a large range of agents and toxins. As many of these agents andtoxins are highly infective, the platform should demonstrate sensitivityand specificity to allow early exposure detection, reduce falsepositives, enable targeted countermeasures, and minimize the spread ofinfection. The platform should also allow for rapid detection to enabletimely intervention. In this context, the challenges of developing asensitive, yet specific, high-throughput detector having a wide workingrange may be appreciated. The challenges are further complicatedconsidering the need for portability, minimal operational complexity,low power consumption, low manufacturing cost, and operability in harshenvironments, for example.

Platforms for detecting molecules have evolved from impractical andlaboratory-based systems to portable miniaturized “Lab-on-a-Chip”platforms. For example, the detection of biological threats has evolvedfrom conducting threat detection and diagnosis though the LaboratoryResponse Network to detection using a mobile lab based system, such asthe Biological Integrated Detection System (BIDS), to mesoscale peptidebioassays. This evolution is representative of the need for smallmolecule detectors that are capable of rapid and point-of-use detection.

Traditional bioassays fall into two categories: label-based orlabel-free. In label-based bioassays, the target molecule, such as atoxin or other molecule, binds with a bait molecule, often acomplementary peptide, DNA, or RNA molecule which has a covalentlyattached label. Fluorescent dyes and radioactive isotopes are commonlyused labels where binding of the target molecule to the bait moleculecauses the release of fluorescence or radiation. In this context, themeasurement of fluorescence or radioactivity provides an indirectdetection and quantification of the target molecule.

However, these array label-based assays suffer from significantlimitations despite some improved sensitivity and specificity. First,these array label-based systems require identification, design,synthesis, and immobilization of the bait molecules, which aresignificantly rate-limiting in the assay manufacturing process. Second,immobilization of a bait molecule with a three-dimensional structureresults in a loss of activity of the bait molecule which may generate afalse negative outcome. Third, the addition of a covalently boundfluorophore or other radioactive tag significantly modifies aninteraction between the target molecule and the bait molecule, resultingin false positives and negatives. Fourth, tagging a bait molecule with afluorescing or radioactive tag adds a layer of complexity to themanufacturing process. Fifth, the assay requires that readers detect theoptical/radiation signal from the tags be incorporated with theplatform, thus dramatically increasing platform cost while reducingportability. Finally, the extinction of a signal generated from abinding event due to scattering from the background matrix is apersistent problem.

In the context outlined above, the limitations imposed by traditionallabel-based bioassays prompted the development of label-free methods.Like the label-based bioassays, a label-free bioassay includes baitmolecules immobilized on a solid substrate. The detection of the bindingbetween the target molecule and bait molecule is based on (a) the changein charge at the solid-liquid interface that results from the bindingevent, (b) evanescent wave attenuation due to a change in refractiveindex at the solid-liquid interface, and/or (c) mass change at thesolid-liquid interface. Charge based detection methods eliminate theneed for expensive signal readers, thereby reducing the cost ofdetection, enhancing system portability, reducing overall powerconsumption, and increasing ease of operation. The charge based methodis also scalable, which is an essential strategy in developing a highthroughput detection platform. Though the label-free platforms do notsuffer from problems like tag-altered target molecule binding andreduced signal yield, they are still afflicted by the issue of baitmolecule misfolding on immobilization to a solid surface.

Generally, the bait molecule is utilized to infer whether the targetmolecule is present or absent in both label-based and label-freeplatforms. The actual identity of the target molecule is inferred fromthe nature of the bait molecule with which binding occurs. Massspectrometry, on the other hand, is a time-critical, broadband analysistechnique that directly measures molecular composition from estimates ofcharge-to-mass ratios of vaporized fragments of the analyte. Commercialmass spectrometers are reportedly capable of detection in the nanomolarconcentration range. Arrayed, multi-channel, modular architectures fortime-of-flight (TOF) mass spectrometers have been detailed for rapid,in-parallel acquisition of information.

However, mass spectrometry analysis is better suited to larger molecularweight target molecules that can be fragmented into several constituentmoieties for analysis. Small molecular weight (<5 kDa) target moleculesare not easily identified by this technique. Mass spectrometer andassociated ancillary equipment (e.g., vacuum pumps) are energy intensivein operation and are not easily miniaturized, thus making portability anissue. Additionally, mass spectrometer operation and data analysisrequire intervention of skilled technicians, making the detectionplatform ill-suited for point-of-use applications. Thus, in view oftraditional detection systems, the need for a robust, rapid, low-cost,point-of-use detection platform for small amounts of molecules in fluidsamples can be appreciated.

Molecular vibration-assisted-charge transfer between an electron sourceand donor has been documented in nature. Fruit flies detect odorants bytransferring an electron from an intracellular electron source uponentrance of an odorant into a transmembrane pocket. The electron chargetransfer stimulates G-protein mediated signal transduction pathways andthus allows the fruit fly to identify an odorant utilizing vibrationalsignatures of odorant molecules. Similarly, according to aspects of theembodiments described herein, the detection of molecular analytes by thedetection of electron transfer is achieved. In the biosensor, accordingto the embodiments described herein, current measured due to electrontransfer that contains information about vibrational frequencies ofmolecular bond vibrations within a molecular analyte, as well asinformation about participating electronic energies, is acquireddirectly from the engineered inorganic transducing interface andanalyzed.

Generally, the biosensor system according to the embodiments describedherein includes an electrochemical charge transfer platform where theapplication of an electromagnetic field at an interface between anelectrode and an electrolyte results in the transfer of charge acrossthe interface. The transfer of charge may occur from the electrode to achemical species in the electrolyte that can accept the charge (i.e., aredox-active species) or vice-versa. The transfer of charge is, in turn,characterized by electromagnetic field-mediated tunneling of electronsthat may be assisted by exchange of energy with thermal vibrations ofother non-redox-active species (i.e., analytes) at the interface. Theinterface is engineered such that a number of collisions experienced bytransferring electronic charge with other analyte molecules is minimalbut not zero. The collisions of the tunneling electrons with thermalvibrations are responsible for the energy exchange between thetransferring charge and the analyte molecules.

The electrochemical charge transfer platform according to theembodiments described herein includes a metal/semiconductor electrodeand an organic or aqueous electrolyte separated by a thin dielectriclayer. The organic or aqueous electrolyte, which is coupled or inimmediate contact with the thin dielectric layer, is characterized by adistribution of uni-polar charge that decays to zero as distance fromthe dielectric-electrolyte interface increases. The dielectric layeracts as a molecular insulator that slows down the rate of electrontransfer sufficiently such that a tunneling electron minimally collideswith surrounding thermal vibrations. Measured current that wouldcharacterize the tunneling of electrons across a suitably engineeredinterface would contain signatures of the resonant energy exchangebetween the tunneling electrons and the molecular vibration modes of theanalytes, as well as signatures of the electronic energies in theelectrode and redox active species that participate in the tunnelingprocess.

The biosensor system according to the embodiments described hereinfurther includes a high gain noise suppression feedback loop toelectronically “cool” the system and minimize thermal noise thatotherwise dissipates the resonant signal of interest. At low electronictemperatures, transfer of electronic charge occurs in a resonant mannerby inelastic interactions with quantized vibrations of a target analyteas well as by direct elastic interactions between the participatingelectronic energy levels.

In various aspects and embodiments, the biosensor system measures atleast one of resonant interactions by measuring a) the tunneling current(I) as a function of applied voltage (V), b) small signal conductance(dI/dV) as a function of applied voltage, or c) conductance derivative(d²I/dV²) as a function of applied voltage. Each resonance featuremanifests as a discontinuity in the measured profiles and may becorrelated to a vibrational frequency of a molecular bond in the analyteor to a participating electronic energy level. Since vibrationalfrequencies may be relied upon as characteristic signatures of molecularbonds, akin to human fingerprints, for example, the number and types ofbonds in the analyte can be determined from these discontinuities.Discontinuities corresponding to electronic energy levels yieldinformation specific to the electronic structures of the electrode andelectrolyte phases that may themselves be perturbed by the analytechemistries. Each analyte possesses a unique molecular bond signature,thus allowing direct, highly specific analyte detection.

With further regard to resonant electron transfer at an electrochemicalinterface, the biosensor system described herein relies in part uponmeasuring electron flux produced in charge-transfer-relatedquantum-mechanical transitions at an electrochemical interface. In thiscontext, the measured electron flux or currents are representative ofmolecular structural and chemical information where quantum mechanicaltransitions manifest as discontinuous features in the currents. Themolecular structural and chemical information, once determined, isunique to each analyte, thus allowing for highly specific molecularspecies determination.

Turning now to the drawings, the features and aspects of the embodimentsare described in further detail.

FIG. 1A illustrates an example schematic diagram of an electrochemicalinterface with characteristic length scales to determine the nature of acharge transfer reaction according to aspects of the embodiments. Chargetransfer across an electrified electrode-electrolyte (orelectrode-insulator-electrolyte) interface may be limited by severalfactors, such as a) mass transport of reactants to theelectrode-electrolyte interface, b) capacitive charging/discharging ofthe electrode-electrolyte interface, or c) quantum-mechanical tunnelingof electrons from electrode to redox energy levels or vice-versa. Whenthe electrochemical interface is engineered such that charge transfer islimited by the electronic transition process, the nature of the electrontransition and the magnitude of the transition charge flux depends onthe extent of electronic coupling between the initial and finalelectronic energy states of the transferring electron as well as thestrength of the nuclear-electrostatic coupling between the electron andthe thermal molecular vibrations.

The strength of the coupling factors can be well represented byequivalent length scales. For differing values of length scaleparameters, the nature of the transition process is qualitativelydepicted in FIG. 1A. Length scale parameters are themselves depicted asfunctions of interface charge density Q(0) and, thus, the nature of thetransition process can be modulated by active control of the interfacecharge density. One optimal charge transfer regime suitable for thetransduction of vibrational mode information from the electron tunnelingprocess relies upon an “intermediate” strength of the two couplingenergies and hence an “intermediate” value of interface charge density,as depicted in FIG. 1B.

The electronic and electronic-nuclear coupling strengths can be tuned inmany different ways, for example, by changing the applied electrostaticfield, by tuning the local interface chemistry, conditioning thephysical system to reduce its intrinsic noise, scaling down the physicalsensor interface, and combinations thereof. In addition, assistingelectromagnetic fields (e.g., optical and magnetic fields) may also berelied upon to induce electronic transitions between energy levels inthe electrode-electrolyte system that are resonant with the dissipatedenergy of the field. Control of the above mentioned parameters reducesthermal de-phasing of the resonance phenomena in the charge transferprocess.

The coupling between electronic energy states participating in thetransition process and the surrounding bath of thermal vibrational modescan be weak, as illustrated by the example in FIG. 2A, or strong, asillustrated by the example in FIG. 2B. Further, coupling strength may betuned by applied bias, interface chemistry, interface size, intrinsicinterface noise, the application of electromagnetic fields, orcombinations thereof, for example. As illustrated in FIG. 2B, when anapplied bias allows for electron transition where electronic energiesare significantly coupled to vibrational modes, de-phasing is strong. Ina strongly coupled electron transfer, the electron wavefunction islocalized to initial and final energy states before and after the chargetransition. As illustrated in FIG. 2B, this results in particle-likebehavior and a thermalized non-adiabatic charge transfer event. On theother hand, in the case of weak coupling between the electronic energystates and the molecular vibrational energy levels, the electronicwavefunction is delocalized over initial and final electronic energystates. As illustrated in FIG. 2A, this results in a wave likeinteraction of the electron with the surrounding vibrational modes andenables the resonant transduction of vibrational mode information.

The coupling between the discrete electronic energy states of theelectrode and the redox-active species in the electrolyte also affectsthe ability of the interface to transduce the molecular vibrational modeinformation. A strong coupling between the electrode-electrolyteenergies results in a “fast” charge transfer event that is limited onlyby the rate of dielectric thermal repolarization around the electrodeand redox-active species, as illustrated by the example in FIG. 2C.Importantly, for this kind of charge transfer reaction, referred to asan “adiabatic” reaction, the transitioning electron is always in aground state resulting in no possibility for resonant electron transferto occur. On the other hand, for the case where the coupling betweenelectrode and electrolyte levels in very weak, there is littleinteraction between the two phases of the system and charge transfer isyet again mediated by thermal excitation only, as illustrated by theexample in FIG. 2D.

An optimum level of electronic-electronic and electronic-nuclearcoupling is required to transduce the discrete vibrational modeinformation as indicated previously. Thus, in the optimal case, theelectron transfer is limited by the rate of the electronic tunnelingtransition from reactant to product state, where the electronparticipates in an inelastic exchange of energy with the molecules inthe intervening layer between the electrode and redox active species inthe electrolyte. This optimally-coupled transition allows thetransferring electron to be de-excited from a higher energy level to alower energy level, thereby losing energy to the intervening molecularspecies, which shows as a signature in the current, conductance, orconductance derivative signal.

With regard to the design of a vibrational mode information transductioninterface, according to aspects of the embodiments described herein, themeasurement of the flux of electrons crossing an electrified dielectricmonolayer modified electrochemical interface allows for analytedetection. In this context, FIG. 3 illustrates an example schematicdepicting the measurement of a flux of electrons crossing an electrifieddielectric monolayer modified electrochemical interface and theinteraction of the tunneling electrons with an analyte co-located at theinterface according to aspects of the embodiments.

In one embodiment described herein, a sensor consists of an electrode(e.g., metal/semiconductor), a molecularly thin spacer layer, and aredox-active species in the electrolyte. An electrode that acts as asource or sink of transitioning electrons may be defined by discreteelectronic energy states that can interact with discrete energy levelsof the molecular redox species in the electrolyte, as opposed to acontinuous collection of energy levels that are characteristic of amacroscopic wire. The need for a discretized energy structure of theelectrode at room temperature tends to the need for an electrode ofnanoscale dimensions. The nanoscale electrode would, in turn, beelectrically addressed by a lead (e.g., electrical lead) that applies orsupplies a suitable voltage and, as a result, charge flows in anexternal instrumentation circuit as a tunneling current.

The sensor size, lead area, dielectric spacer thickness, choice ofelectrolyte (e.g., aqueous, organic, ionic salt), and choice ofredox-active species in the electrolyte may be determined, for example,so as to optimize the electronic and electronic-nuclear coupling at theelectrochemical interface. Quantitatively, “optimal” is defined in thiscontext by a specific value of interface charge density. This value ofinterface charge density may be determined by kinetics of theaccompanying electron transfer reaction (which determines the nature ofelectrode material, the nature of electrolyte, and type of redox activeion in the electrolyte) and dielectric spacer thickness. Thedetermination of equivalent or suitable lead area is a trade-off betweenminimizing parasitic capacitance from insulated leads and electricaldouble layer at the solid-liquid interface and minimizing the thermalbroadening of the discrete electronic energies of the nano-electrodewith increasing size.

FIG. 4 illustrates design considerations for a transducing interface andfactors in view of quantum-mechanical to classical transition behavior.As shown in FIG. 4, the total sensing area as well as the sensing tolead area ratio may be designed in view of the quantum-mechanical toclassical transition behavior of the system as well as the transitionbetween weak and strong nuclear-electronic coupling regimes.Additionally, a constraint on the upper value of the lead area isdetermined by estimating the extent of thermal broadening induced by amacroscopic lead that electrically addresses the nanoscale electrode. Insome embodiments, an intervening molecularly insulating spacer may beutilized to weakly couple the macroscopic lead to the nanoscaleelectrode. At least in part, the choice of spacer material anddimensions and the total lead area determines or bears upon theeffective broadening of the electronic energies of an electron in thenanoscale electrode. Thus, a suitable mix of these parameters may bechosen among embodiments to ensure that thermal broadening is below thethermal energy at room temperature (˜25 meV). The electron flux ortunneling current at this nano-structured interface is measured eitherdirectly as a current or as an impedance/derivative of system impedancewith applied voltage. In this context, by the application of suitabledata analysis techniques, detailed structural information about amolecular analyte can be obtained.

With the application of a voltage between a macroscopic lead and areference electrode that sits in bulk electrolyte solution, electronstunnel from the nanoscale electrode to the redox-active species in theelectrolyte. If the interface is engineered appropriately, for “optimal”coupling conditions, such that the tunneling of the electron fromelectrode to electrolyte is rate limiting and no other process (e.g.,mass transfer of redox-active species to interface from bulkelectrolyte, capacitive charging/discharging of interface charge, ortunneling of electrons from lead to nanoscale electrode) is slow enoughto compete, then a current measured by a low noise transimpedanceamplifier and acquired by a data acquisition system corresponds to adirect measurement of this tunneling event.

In other words, as an electron tunnels across the appropriatelyengineered interface, it loses energy equivalent to the applied biasvalue, and this energy is lost to molecular vibrations of analytespecies with suitable vibrational energies that exist at the interfacebetween the electrode and redox-active species. Thus, the biosensoraccording to the embodiments described herein measures a spectrum ofmolecular vibrational oscillation modes of an analyte at anelectrochemical interface within a liquid electrolyte in resonance withan energy gap between initial and final electronic energy states of theelectrochemical interface. In addition to vibrational signatures, thetunneling electron also transduces information about electronicresonances arising from elastic (i.e., collision-less) transitionsbetween the electrode and redox species energy levels. However, it isexpected that elastic transitions would be probabilistically less likelyfor a suitably designed interface.

In one sense, this approach is analogous to that of electromagneticprobes, such as near-infrared (NIR) vibrational spectroscopy probes,than with conventional electrostatic measurements. However, as themolecular structure information is transduced directly to an electronicsignal before acquisition, the proposed biosensor is highly scalable.The direct acquisition of chemistry specific information about ananalyte in the form of molecular vibrational modes also eliminates theneed for time and labor-intensive combinatorial screening againstbait-molecule probes required by traditional bioassays.

The quantum information transduction mechanism achieved according to theembodiments described herein enables highly specific interrogation ofTHz frequency molecular vibrations at experimentally accessible (˜mV)electronic energies/potentials by scanning the electronic energy with anapplied voltage at a metallic, electrically conductive lead.Experimental results, such as those illustrated in FIG. 5, suggestsensitivity of the biosensor described herein to a single atom massisotope substitution, as well as sensitivity to structural isomerism,which has not been demonstrated before with traditional electronicdetection techniques.

According to aspects of the embodiments, various types of biosensorstructures and interfaces may be relied upon to specifically optimizeelectronic and/or electronic-nuclear coupling. For example, various thin(e.g., sub˜1 nm) dielectric-film-modified nanoscaleelectrode-electrolyte interfaces may be relied upon. The interfaces maybe patterned in planar fashion on a silicon die using standard planarmicrofabrication techniques, for example. FIGS. 6A-C illustratedifferent types and structures of biosensor interfaces. Depending uponthe type of the interface, a liquid or other sample may either bepositioned upon or over the interface. Alternatively, the interface maybe inserted or immersed in the sample.

Depending upon the type of the biosensor interface, one or moreelectrodes may be planar with metallic rectangular pads being used forcontacts and thin leads being used for the sensing architecture. Tocontrol the volume of fluid, a liquid fluidic channel/chamber may beused to contain the volume of liquid sitting atop thin leads of thebiosensor interface. In some embodiments, the entire biosensor interfaceelectrode structure may be fabricated on a silicon substrate usingstandard microfabrication techniques, and the fluidic channel can bemade out of a plastic or ceramic and sealed hermetically with thesilicon surface to create a leakproof system.

Turning to FIG. 6A, one example of an electrode-electrolyte interface isillustrated. The interface is designed to specifically optimizeelectronic and electronic-nuclear coupling. In another embodimentillustrated in the example of FIG. 6B, thenanoscale-electrode-dielectric film-electrolyte interface can belocalized at the tip of a sharpened probe which may then be insertedinto a volume of interest to characterize the spatiotemporal chemistryof the local environment. For the embodiment of FIG. 6B, the tipstructure may be fabricated out of an insulator such as glass orplastic, and a thin metal lead is extended to the end of the tip where asensing electrode exists. In yet another example illustrated in FIG. 6C,leads to electrically address a nano-engineered interface are designedto be “through substrate” rather than planar as mentioned in the firstscheme. In this configuration, the substrate is selected to beinsulating, like glass, and the design includes aspects of the first andsecond schemes. It should be appreciated that the example interfacesillustrated among FIGS. 6A-6C are provided by way of example only, andother forms, shapes, and styles of interfaces are within the scope ofthe embodiments.

According to other aspects of the embodiments described below, using oneof the interfaces illustrated in FIGS. 6A-6C, for example, tunnelingcurrent flux is recorded by ultra-low noise acquisition circuitryfabricated, for example, by a complementary metal oxide semiconductor(CMOS) process and integrated with the sensing interface usingheterogeneous integration or other suitable techniques. According toother aspects of the embodiments, shielding and interconnectiontopologies are designed to minimize signal contamination caused, forexample, by band-limited white noise, electromagnetic interferingsignals, flicker noise, and artifacts arising from the digital dataacquisition system. Acquired data may be transmitted off-line forfurther filtering, if necessary, as well as for data recording anddisplay. In certain embodiments, the biosensor further includes meansfor pre-screening a level of specific biological markers before assayingfor an analyte of interest. For example, in a biosensor targeting bloodtoxins, pre-screening for cytokines allows for evaluation of overallhealth and can indicate presence or absence of a bacterial infection.

FIG. 7 illustrates an example biosensor platform 700 according tocertain aspects of the embodiments described herein. Among otherelements, the biosensor platform 700 includes a fluidic system within apackage 730 and a sensor 714. In one embodiment, the fluidic systemincludes an acquisition zone 702 and one or more disposable modules. Thepackage 730 may allow for easy access to and replacement of thedisposable modules. The disposable modules may include a filtrationmembrane 706, an immunoseparation membrane 708, a micro-chromatographcolumn 710, and an absorption pad 712, for example, as illustrated inFIG. 7.

In the biosensor platform 700, the sensor 714 may include anelectrochemical or patterned electrochemical interface and an interfacechip integrated into a low-cost, disposable, lateral flow-basedmicrofluidic architecture. In one example operation, capillary transportmay be relied upon in the biosensor platform 700 to separate serum fromwhole blood and deliver it to an electrode surface of the sensor 714.However, the mechanism to induce fluid flow in the device is not limitedto capillary transport or flow. Dielectrophoresis may also be employedto actuate the liquid medium in the portable biochip configuration.

Among other elements, the sensor 714 may include a plurality of thinfilms 740 (e.g., the electrochemical or patterned electrochemicalinterface), a semiconductor die 750, and an application-specificintegrated circuit (ASIC) 760. The thin films 740 may be deposited byatomic layer deposition, for example, and include working 742, counter744, and reference 746 films or areas. The semiconductor die 750 mayinclude an electrode array and through-die vias for electrical couplingwith the ASIC 760. The ASIC 760 may include bonding pads 762 and 764.The bonding pads 762 may be relied upon for electrical connection withthe through-die vias from the semiconductor die 750, and the bondingpads 764 may be relied upon for electrical connection to otherprocessing and/or data collection processors or circuitry. It should beappreciated, however, that the structure of the sensor 714 illustratedin FIG. 7 is provided by way of example only, as other equivalentstructures are within the scope of the embodiments.

In one embodiment, the biosensor platform 700 includes elements at themacro-, micro-, and nano-scales, where the microfluidic elements bridgethe nano-scale transducer to blood sampling and dispensing at themacro-scale. Since the patterned sensor interface with the integratedelectronic is relatively costly, the microfluidics may be designed suchthat fabrication costs are relatively low, power consumption isnegligible, and the microfluidic component can be easily disposed of ifexcessive blockage obstructs the flow path.

Referring back to FIG. 7 for a description of the operation of thebiosensor platform 700, a sample 704 may be dispersed (e.g., dropped) inthe acquisition zone 702. The sample 704 is then either actively (e.g.,via dielectrophoresis) or passively (e.g., via capillary action) pumpedthrough the fluidic system. In the example platform in FIG. 7, thesample 704 is first wicked through a filtration membrane 706. In oneembodiment, the filtration membrane 706 possesses a graded porestructure capable of separating serum from whole blood. Next the serumpasses through the immunoseparation membrane 708, such as anitrocellulose membrane or other appropriate type of membrane comprisingsurface antibodies specific to high abundance proteins, which remove thehigh-abundance proteins. Finally, the liquid sample moves though themicro-chromatograph column 710, thus fractionating the remainingproteins and results in size separated elutants at the exit of thecolumn 710. In one example embodiment, the micro-chromatograph column710 is comprised of a tapered microfluidic channel containing aphoto-polymerized gel.

It is noted that, although not required for all sample types, thefluidic system is preferred when analyzing complex mediums, such asblood, where components may interfere with the detection of lowabundance analytes. Pumping of the sample 704 may be active or passiveinto the fluidic system. It should be appreciated that the filtrationmedia, chosen filter membranes, other membranes, and characteristics ofthe micro-chromatograph column 710 may be dependent upon factors such assample type, sample amount, or abundance of target analyte, for example.

Continuing with the operation of the biosensor platform 700 in FIG. 7,after the sample 704 passes through the fluidic system, it is exposed toan active interface at the sensor 714. The active interface includes atransducing electrochemical interface integrated with underlyingacquisition electronics, as described herein. An area 716 of the sensor714 is patterned as electrically accessible, thermally insulated,pixilated electrodes for interrogating the sample 704. It should beappreciated that pixel electrodes 718 of the sensor 714 may be singularor exist as an array in a configuration with common counter andreference electrodes. At 720, FIG. 7 further illustrates a magnifiedcross section of view of the transducing electrochemical interface witha sample for evaluation disposed thereon.

FIG. 8 illustrates an enlarged view of sensor 714 with electrode sensorsarranged in an array. The sensor 714 in this embodiment exists as alayered, heterogeneously integrated sensor platform. In one embodiment,the sensor 714 includes a fluidic chamber transfer structure 802 thattransfers a liquid sample from the fluidic system to the electrodesensor array 804. The sample then reaches the electrochemical interfaceelectrode array of the sensor 714, which includes dielectric thin filmsdeposited by self-assembly techniques, atomic layer deposition, and/ormolecular vapor deposition (ALD, MVD). The circuitry 806 of the ASIC 760may be formed using any suitable semiconductor process including, forexample, a CMOS process. The structure and purpose of the circuitry 806is described in further detail below with reference to FIG. 14.

Referring again to FIG. 7, the biosensor platform 700 may be mounted ona shielded, printed circuit board (PCB) for electrical access. Paralleldata acquisition over a large applied bias range is made possible byelectronic-energy-window specific optimization of individual electrodepixels, with each electrode pixel nanostructure being optimized forinterrogating a specific electronic energy/bias window. If necessary,data acquired as a transition current signal can be transmitted to anexternal system for post-acquisition processing, storage, and display.The biosensor platform 700 is designed such that sensing and dataacquisition modules can be easily added or removed so that the platformcan be dissembled, interchanged, and disposed of.

Resolved spectral information, once acquired, is then correlated withvibrational energy data to identify specific molecular speciesassociated with the macro-molecule analyte. This may be accomplished byemploying an information-driven strategy for targeted, non-redundantanalysis of a bio-analyte in an electrolyte solution. Signatures ofinformation-rich subsets of the bio-analyte, such as cysteine-containingpeptides, phosphorylated peptides, or glycosylated peptides, may betracked in the resolved spectrum of the bio-analyte. These subsets willserve as molecular markers for identifying and quantifying the presenceof molecular species of interest. A reference database containing thesemolecular markers may be constructed for each target analyte as furtherdescribed below. In other words, each analyte may be expected to produceits own signature spectrum of information. By comparing the resolvedspectrum from a sample to the reference database, the target analyte maybe identified.

Turning to FIG. 9, processes of spectral data collection 910, referencedatabase collection 930, and analysis 950 are described according toaspects of the embodiments. With regard to the process of spectral datacollection 910, an example of the process 910 is described below inconnection with a sample of raw blood. It should be appreciated,however, that the process 910 may be applied to other types of samples.Further, the process 910 is described below in connection with thebiosensor platform 700 of FIG. 7. Again, it should be appreciated thatthe process 910 may be performed in connection with other biosensorplatforms similar to the biosensor platform 700.

Briefly, among other steps, the process of spectral data collection 910includes pumping a sample through a fluidic system at reference numeral912, filtering the sample at reference numeral 914, separating andremoving at least one composition from the sample at reference numeral916, fractionating the sample at reference numeral 918, and transducinginformation from the sample at reference numeral 920. The pumping,filtering, separating/removing, and fractionating, at reference numerals912, 914, 916, 918, respectively, may be performed in connection withone or more of the disposable modules of the fluidic system describedabove with reference to FIG. 7. Further, the transducing information atreference numeral 920 may be performed in connection with the sensor 714described above with reference to FIG. 7.

As for the more particular example of conducting the process of spectraldata collection 910 using a sample of raw blood, after pumping atreference numeral 912, the sample of raw blood may be subject tofiltering at reference numeral 914, where serum is separated from wholeblood. Next, the serum is cleaned of high abundance proteins atreference numeral 916 by passing through an immunoseparation membrane,such as a nitrocellulose membrane that comprises surface antibodiesspecific to the high-abundance proteins, for example. The liquid sampleis then fractionated reference numeral 918, such that different proteinsfractions are eluted sequentially onto the active sensor area. Theseparation of proteins may occur by utilizing a general protein specificproperty, such as charge-to-mass ratio, to sequentially elutelow-abundance proteins. Finally, information is transduced from theeluted proteins at reference numeral 920.

FIG. 9 further illustrates the process of reference database collectionor generation 930. As one example embodiment of the process 930, atreference numeral 932, the process 930 includes digesting a purifiedrecombinant form of a target molecule (or biological surrogate, in thecase of neurotoxins). That is, the target molecule is systematicallydigested by enzymes, such as trypsin and chymotripsin, to generatepeptide fragments. At reference numeral 934, the digesting is followedby separating using a multi-dimensional separation technique, such as a2-D poly acrylamide gel electrophoresis (2-D PAGE) process in tandemwith high performance liquid chromatography (HPLC, preferably reversephase-HPLC), for example. At reference numeral 936, the process 930includes collecting fractions. That is, fractions are collected,purified, and re-extracted in a suitable buffer and analyzed using thedisclosed quantum tunneling electronic biosensor at reference numeral942. The data, after background subtraction, is analyzed forcharacteristic spectra of moieties specific to the peptide fragment inthe aliquot being tested.

In other aspects of the process 930, stable isotopes of referencepeptides, for example, may also be prepared at reference numeral 938. Insome embodiments, the same fractions as well as isotope-labeledreference peptides prepared at reference numeral 938 may also beexamined in parallel by traditional liquid chromatography-massspectrometry (HPLC-MS-MS) techniques at reference numeral 940.

The analysis process 950 may include, at reference numeral 952, one ormore of acquiring, post processing, and/or displaying data collected bythe spectral data collection process 910. At reference numeral 954, theanalysis process 950 may also include comparing signatures from the datacollected by the spectral data collection process 910 with a database ofreference signatures (i.e., 956) collected by the reference databasecollection process 930. At reference numeral 954, the database ofreference signatures 956 may be compared with raw data from thebiosensor platform 700 to identify molecular analytes of interest. Itshould be appreciated that the processes 910, 930, and 950 are providedby way of example only.

Turning to FIG. 10, example tunneling barriers at metal-dielectric anddielectric-electrolyte interfaces are further described according toaspects of the embodiments. In some embodiments, the tunneling barrierscan be engineered utilizing thin films, such as those included in thesensor 714 of FIG. 4, of organic or inorganic materials with suitableproperties to minimize the electronic coupling between the energy levelsof the electrode and redox active species in the electrolyte. FIG. 10depicts tunneling barriers ϕ₁ and ϕ₂ at the metal-dielectric 1002 anddielectric-electrolyte 1004 interfaces. Coupling between initial andfinal electronic energy states in the optimally coupled electronictransition is modulated by an electrostatic tunneling barrier ϕ₂ locatedat the dielectric-electrolyte interface 1004 as well as by an effectivebarrier limiting charge injection at the electrode-dielectric interface1002. Nanoscale engineering of barrier heights at theelectrode-dielectric thin-film interface 1002 as well as at thedielectric thin-film-electrolyte interface 1004 is utilized to minimizede-phasing of resonant signatures of energy exchange in the electronictransition due to strong electronic coupling.

The desired minimization may be achieved by increasing theelectron-tunneling barrier ϕ₂ at the dielectric-electrolyte interface.In one embodiment, the electron tunneling barrier ϕ₂ increases as theelectrolyte pH increases. Other exemplary embodiments achieve anincreased tunneling barrier by increasing electrolyte anionelectronegativity, increasing dielectric monolayer functional groupelectronegativity, and/or increasing dielectric monolayer thickness, forexample.

The desired minimization in electronic coupling may also be attained byincreasing the limiting barrier φ₂ at the dielectric-electrolyteinterface 1004. In one embodiment, this is achieved by coating thedielectric monolayer 1003 with an organic coating, such as short chainsilanes, with different electronegative, electrolyte-facing functionalgroups, such as —OH, —OR, —COOH, —SH, —SR, —COR, —NO2, —Br, or the like.These aforementioned coatings are suitable for forming with MVD at thedielectric surface 1003.

The metal electrode-dielectric barrier 1002, unlike thedielectric-electrolyte barrier 1004, is a function of the metalwork-function, dielectric band gap, and nature of molecular orbitaldistortion induced by a bond between the metal 1001 and dielectricmaterials 1003. A reduction in the coupling between electrode 1001 andelectrolyte 1005 may also be achieved by altering the tunneling barrierlocated at the dielectric-electrode interface. For example, increasingthe tunneling barrier cpi at the dielectric-electrode surface reducescoupling between electrode and electrolyte phases, thus leading toincreased resolution of vibrational frequency information in themeasured current.

For some embodiments, the mechanism of tunneling based charge injectionin the dielectric 1003 would be electron tunneling. In other words, thedielectric 1003 would be comprised of an inorganic-oxide. For theseembodiments, metals such as Pt, Ir, Se, or Au, or their alloys indifferent compositions are preferred. For other embodiments,hole-tunneling is the mechanism of charge injection in the dielectric1003. It should be appreciated that the dielectric or dielectric film1003 in these embodiments may be comprised of an organic alkane. Forthese embodiments, metals like Ta, Ti, Zr, Hf, or their alloys invarious compositions may be preferred. The final metal choice isdependent on many factors including mechanical, diffusional, andelectrochemical stability of the electrode, ease of deposition,electrical resistivity, and ability to seed a dielectric layer, forexample.

Among embodiments, nanoscale structures of metal electrodes for sensorsdescribed herein are fabricated either in top-down methods usingnanoscale patterning techniques like Electron Beam Lithography (EIB) orFocused Ion Beam (FIB), or bottom-up methods like nanoparticleself-assembly on patterned structures or using a combination of methodsthereof.

In various embodiments, the dielectric film 1003 that spatiallyseparates the electrode 1001 and electrolyte 1005 layers is comprised ofa medium-k nanolaminate. The high-k material in this nanolaminate may beTa₂O₂, ZrO₂, TiO₂ or other suitable material. It is noted that largedielectric constants for the insulating film facilitate greater chargeaccumulation at the dielectric-electrolyte interface, thus effectivelyincreasing the tunneling barrier and reducing the electronic coupling.However, the increased charge density increases the nuclear-electroniccoupling. Also, since a larger dielectric constant is typicallyassociated with small band-gap and, consequently, higher non-tunnelingleakage current, the high-k material may be intercalated betweenalternating layers of lower-dielectric constant oxides.

In the context outlined above, FIG. 11 illustrates sequential layeringof high 1102 and low 1104 k-dielectric materials for a high-k dielectricinsulator according to aspects of the embodiments. The final choice ofmaterials used to form a given dielectric nanolaminate may be determinedby insulator properties like breakdown resistance, electrochemical andmechanical stability, and chemical inertness to aqueous electrolytes inthe presence of an applied bias. Other low-k materials, like molecularorganic spacers (e.g., derivatized alkane/alkene thiols and derivatizedsilanes) may also be used as functional insulating spacers for thenanoscale interface, for example.

On the other hand, the dielectric film utilized to insulate theaddressing lead from the electrolyte solution would typically be oflow-k material like SiO2. The low-k nature of the insulating dielectricwould minimize losses induced by the parasitic capacitance thatcontributes to the dephasing of the resonance signal. The sensorinterface configuration may thus be comprised of low-k and high-kinsulating material co-patterned on the same interface depending on thefunctional utility of the insulator.

Reduction in nuclear-electronic coupling may, additionally oralternatively, be achieved by applying a directional magnetic field tothe nanoscale electrochemical interface, where a) the dielectric filmthat serves as the function insulating element includes a nanolaminatestructure that uses differentially oriented film magnetic moments toconstrain the spin of the tunneling electron, and b) the nanoscaleelectrode participating in the redox reaction would comprise ananostructured ferromagnetic/paramagnetic element with strongly orientedelectronic moments. Preferably, the magnetic tunneling nanolaminate willcomprise dielectric-based thin-film architectures with room temperatureferromagnetic properties that allow for the generation of local,inhomogeneous magnetic fields that can interact with the magnetic dipoleof the transitioning electron to “gate” the quantum-mechanicaltransition.

In the context outlined above, FIG. 12 illustrates a magnetic tunnelingfilm architecture that uses differentially oriented film magneticmoments to further restrict electronic transition according to aspectsof the embodiments. A multi-stage-gate-like design of the dielectricnanolaminate may enable further minimization of the de-phasing resultingfrom nuclear-electronic coupling between the energy of the tunneling andthe surrounding bath of thermal vibrations. In one embodiment, thedielectric film includes low-k (e.g., HfO₂) 1204/high-k 1202 dielectricsubstacks interspersed with thin films of a non-magnetic dielectricinsulator 1206 with an intermediate value of dielectric permittivity,such as Al2O3. Aluminum in the Al₂O₃ thin-film lack the d-shell orbitalsnecessary for displaying magnetic susceptibility and thus the aluminathin-films are believed to be non-magnetic, making them suited for thisapplication. In other words, the functional ferromagnetic elements ofthe dielectric thin film are made up of low-k (e.g., HfO₂) 1204/high-k1202 dielectric substacks and every two subs-stacks are separated byAl₂O₃ thin-film which functions as an insulating barrier that minimizesdissipative magnetic coupling between adjoining ferromagneticsub-stacks. The total number of substacks and Al₂O₃ thin-films and,thus, the extent of ferromagnetic-induced decoupling, is limited by theoverall thickness of the dielectric film, which is less thanapproximately 10 nm in various embodiments.

Further reduction in nuclear-electronic coupling may be made possible bythe application of a “reaction” gate to control the interface chargedensity at the electrochemical interface. The electrochemical reactiongate proposed in this embodiment would utilize a fast/adiabaticelectrochemical electron transfer reaction to set the charge densityacross the entire solid-liquid interface, which would also include asmall sensing interface area, as illustrated in the example of FIG. 13A.The interface charge control mechanism in this case is referred to as anelectrochemical reaction gate because, like a traditional metal oxidesemiconductor (MOS) gate structure, the reaction gate uses an appliedvoltage (independent of sensor biasing voltage) to control the chargedensity at the interface. However, unlike a traditional MOS gate, thereaction gate relies on a fast electrochemical charge transfer reactionto set the surface charge density to the desired optimal value to ensurean optimal level of electronic and electronic-nuclear coupling. Thisembodiment of the sensor interface may be advantageous in that itdecouples the modulation of the interface charge density from thesensing function, and facilitates an independent modulation of thecoupling mechanisms that can dephase the electronic resonances. Thedimensions of the sensing and gate electrode regions may be settled uponbased on the design rules described herein. In this embodiment, the gatevoltage and the voltage applied to cause the sensing electronictransition are independent. However, the electrochemical reactioncausing the gating effect utilizes the same redox-active species in theelectrolyte as the sensing interface.

In yet another example embodiment of the gate-electrode system, asillustrated in FIG. 13B, the interface is configured to ensure that thegate control is substantially or completely decoupled from the sensingfunction, where the electrolyte and redox active species responsible forthe fast/adiabatic reaction that sets the interface charge density areencapsulated in a fabricated nanofluidic structure that reduces the massand ion exchange with the sensing electrolyte containing the analyte tobe sensed. The nanofluidic structure, however, permits electricalconnectivity between the “gating electrolyte” and “sensing electrolyte”and, therefore, a uniform or near-uniform interface charge density maybe set across the entire solid liquid interface. The nanofluidicstructure described allows for the inter-diffusion of small ionicspecies like protons and hydroxyl radicals between the gatingelectrolyte and the sensing electrolyte, while sterically blocking andpreventing contamination by the larger redox active ions.

A third embodiment of the gate-controlled sensing interface wouldconsist of the sensing interface being localized on the tip of asharpened probe, with the gate electrode being co-located on the body ofthe probe. As illustrated in the example of FIG. 13C, the gate andsensing regions of the device may be segregated by a nanoporous fritfilter that allows for electrical communication between the gatingelectrolyte and the sensing electrolyte.

Another key component to the biosensor platform 700 (FIG. 7) is theanalog front-end instrumentation, which is relied upon for conditioningvoltage signals to measure vibrational frequency signatures in theelectrochemical tunneling current. The conditioning of the appliedvoltage signals involves utilizing a relatively large gain feedback loopto minimize and set the noise present at the electrochemical interfaceto a predetermined value. The noise at the electrochemical interface maybe due to physical or electronic sources present either at the interfaceor in the measurement instrumentation respectively. The physical sourcesof noise are correlated positively with the parameters that promoteincreased electronic and electronic-nuclear coupling and therefore,manipulation of physical noise at the electrochemical interface enablesdirect and active control of the coupling effects. By contrast,engineering of the physical interface, as described herein, is a passivecontrol on the coupling mechanisms.

Besides physical sources of noise, electronic sources of noise from themeasurement instrumentation (e.g., wide band thermal noise, widebandshot noise, 1/fα noise, etc.) also manifest across the electrochemicalinterface and may manifest as enhanced coupling. Therefore, electronicinstrumentation should be designed to minimize electronic noise and toset the physical noise to pre-determined, desired levels.

The measured non-adiabatic current is a function of two non-interactingfrequency domains: (a) a “macro-frequency” (approximately 1 Hz) thatdetermines the rate-limiting step in the macroscopic electrochemicalsystem and (b) a “micro frequency” (>10¹² Hz) that measures the dynamicsof molecular vibrations and the tunneling process, where the dynamicsare manifest in the electronic energy (or equivalently, the appliedbias) space. Multiple measurement schemes involving different voltagesignal types and differing forms of data acquisition are proposed forthe identification of signatures in the measured current.

In one case, a small amplitude low-frequency Alternating Current (AC)voltage excitation is combined with a Direct Current (DC) voltage biasand applied to the electrochemical interface. Then, the lock-in acquiredAC current is recorded as a function of Direct Current (DC) bias at thenano-engineered electrode-electrolyte interface. In another case, a DCvoltage is applied directly to interface and a DC current is acquired.In still another case, a small amplitude low frequency AC voltage iscombined with a DC bias and applied at the interface and higherharmonics of the measured current are acquired with lock-in techniques.The application of the DC and AC voltages and simultaneous acquisitionof the current may be accompanied by the automated modulation of appliedmagnetic fields or noise power set-points. Effective signal extractionrequires suppression of extrinsic noise contributions and control ofintrinsic noise contributions. The acquisition of the tunneling currentsignal (AC and DC components) requires implementation of suitablehardware and software-based data filtration techniques to minimizeelectronic noise picked up in the measurement process.

Turning to FIG. 14, an example of three-electrode feedback suppressionof thermal noise for electronic transition measurements is illustratedaccording to aspects of the embodiments. In FIG. 14, a three-electrodeanalog measurement topology is used with high gain feedback incorporatedto suppress the voltage-noise in the macro-frequency domain. Current isacquired with a transimpedance amplifier topology, where the tunnelingcurrent flowing across a resistor is converted into a voltage that ismeasured. Set values of passive components in the transimpedanceamplifier signal chain allow for the measurement of AC or DC currents.Within this embodiment, the instrumentation system can apply a low noisevoltage (DC, DC+AC) and measure AC or DC currents. Real and imaginarycomponents (i.e., resistive and capacitive current contributions,respectively) of the measured AC current or higher harmonics of thesignal can be isolated using traditional lock-in detection techniques.

Thus, with reference to FIG. 14, a three-electrode analog measurementtopology circuit 1400 may be used for high gain feedback suppression ofvoltage-noise in the macro-frequency domain. The circuit 1400 includes afirst gain amplifier 1402, a second gain and summing amplifier 1404coupled to an output of the first gain amplifier 1402, and a counterelectrode 1406 coupled to an output of the second gain amplifier 1404.The circuit 1400 further includes a reference electrode 1408 and aworking electrode 1410, each coupled to a difference amplifier 1412. Anoutput of the difference amplifier 1412 is coupled to a DC generator1414. Outputs of the DC generator 1414 and an AC generator 1416 arecombined at a node 1417 in the circuit 1400 and provided as an input tothe second gain amplifier 1404. As illustrated in FIG. 14, the combinedoutputs of the DC and AC generators 1414 and 1416 are further combinedwith the output of the first gain amplifier 1402 at the node 1417 beforebeing provided as input to the second gain amplifier 1404.

The circuit 1400 further includes a current measurement circuit 1418coupled to the working electrode 1410 and a lock-in detection circuit1420 coupled to an output of the current measurement circuit 1418. Asillustrated in FIG. 14, a reference signal output of the lock-indetection circuit 1420 is provided as an input to the AC generator 1416,an out-of-phase signal output of the lock-in detection circuit 1420 isprovided as an input to the first gain amplifier 1402, and an in-phasesignal output of the lock-in detection circuit 1420 is provided as aninput to a data acquisition element 1422 for processing.

Although embodiments have been described herein in detail, thedescriptions are by way of example. The features of the embodimentsdescribed herein are representative and, in alternative embodiments,certain features and elements may be added or omitted. Additionally,modifications to aspects of the embodiments described herein may be madeby those skilled in the art without departing from the spirit and scopeof the present invention defined in the following claims, the scope ofwhich are to be accorded the broadest interpretation so as to encompassmodifications and equivalent structures.

Further, it should be noted that ratios, concentrations, amounts, andother numerical data may be expressed herein in a range format. It is tobe understood that such a range format is used for convenience andbrevity, and thus, should be interpreted in a flexible manner to includenot only the numerical values explicitly recited as the limits of therange, but also to include all the individual numerical values orsub-ranges encompassed within that range as if each numerical value andsub-range is explicitly recited. To illustrate, a concentration range of“about 0.1% to about 5%” should be interpreted to include not only theexplicitly recited concentration of about 0.1% to about 5%, but alsoindividual concentrations (e.g., 1%, 2%, 3%, and 4%) and the sub-ranges(e.g., 0.5%, 1.1%, 2.2%, 3.3%, and 4.4%) within the indicated range. Theterm “about” can include traditional rounding according to significantfigures of numerical values. In addition, the phrase “about ‘x’ to ‘y’”includes “about ‘x’ to about ‘y’”.

1-20. (canceled)
 21. A sensor comprising: a biosensor interfaceincluding an electrode and a dielectric film deposited on the electrode,wherein the biosensor interface is configured to operatively couple witha sample comprising a redox specie and an analyte specie; and aprocessing logic operatively coupled to the electrode, and configured toapply a voltage bias between the sample and the electrode, the appliedvoltage bias configured to generate a tunneling current configured toflow between the redox specie and the electrode via the dielectric film,wherein the tunneling current is indicative of the analyte specie. 22.The sensor of claim 21, further comprising a probe including a tip at adistal end of the probe, wherein the biosensor interface is coupled tothe tip of the probe.
 23. The sensor of claim 22, further comprising: asecond biosensor interface including a second electrode and a seconddielectric film deposited on the second electrode, wherein the secondbiosensor interface is configured to operatively couple with the sample,wherein the processing logic is operatively coupled to the secondelectrode, and configured to apply a second voltage bias between thesample and the second electrode, the applied second voltage biasconfigured to generate a second tunneling current configured to flowfrom the redox specie to the second electrode via the second dielectricfilm, wherein the second tunneling current is indicative of the analytespecie.
 24. The sensor of claim 23, further comprising a second probeincluding a tip at a distal end of the second probe, wherein the secondbiosensor interface is coupled to the tip of the second probe.
 25. Thesensor of claim 22, further comprising a fluidic system enclosing avolume configured to receive the sample.
 26. The sensor of claim 25,wherein the enclosed volume of the fluidic system is configured toreceive the distal end of the probe, wherein the biosensor interfacecoupled to the tip of the probe is configured to contact a volume of thesample in the fluidic system.
 27. The sensor of claim 25, wherein thetunneling current is indicative of the analyte specie in the samplevolume.
 28. The sensor of claim 22, wherein the tip is an insulator. 29.The sensor of claim 21, wherein the dielectric film comprises asequential layering of low-k and high-k dielectric materials.
 30. Thesensor of claim 29, wherein a layer of low-k material and a layer ofhigh-k material are separated by a layer of non-magnetic dielectricinsulator.
 31. The sensor of claim 29, wherein the high-k dielectricmaterials includes one or more of Ta2O2, ZrO2 and TiO2.
 32. The sensorof claim 21, wherein the processing logic is coupled to the electrode bythrough-silicon vias.
 33. The sensor of claim 21, wherein the processinglogic includes a voltage source configured to apply the voltage bias,and the applied voltage bias is further configured to produce aweakly-coupled non-adiabatic electron flux across the dielectric film.34. The sensor of claim 21, wherein the processing logic includes a lownoise transimpedance amplifier configured to detect the tunnelingcurrent.
 35. The sensor of claim 34, wherein the detected tunnelingcurrent is indicative of molecular vibrational states of the analytespecie located at the biosensor interface.
 36. A method comprising:detecting, by a sensor, a signature tunneling current indicative of oneor more vibrational states of a first analyte in a sample comprising aredox specie and the first analyte, wherein the sensor is configured toapply a voltage bias between the sample and the sensor, and generate thesignature tunneling current between the redox specie and the sensor;comparing data characterizing the detected signature tunneling currentwith data characterizing a plurality of signature tunneling currentsassociated with one or more analytes; and determining an identity of thefirst analyte based on the comparison.
 37. The method of claim 36,wherein the sensor includes: a biosensor interface including anelectrode and a dielectric film deposited on the electrode, wherein thebiosensor interface is configured to operatively couple with a firstportion of the sample; and a processing logic operatively coupled to theelectrode, and configured to apply the voltage bias between the firstportion of the sample and the electrode.
 38. The method of claim 37,wherein detecting the signature tunneling current further comprises:filtering the sample to separate the first portion of the sample from asecond portion of the sample, the first portion includes the firstanalyte and a second analyte; passing the first portion of the samplethrough an immunoseparation membrane; and directing the first analyteand the second analyte to the sensor, wherein the first analyte arrivesat the sensor during a first time duration and the second analytearrives at the sensor during a second time duration.
 39. The method ofclaim 38, further comprising detecting by the sensor, a second tunnelingcurrent indicative of the second analyte during the second timeduration.
 40. The method of claim 38, wherein the immunoseparationmembrane includes surface antibodies.
 41. The method of claim 38,wherein the immunoseparation membrane includes a nitrocellulosemembrane.
 42. The method of claim 38, wherein the first analyte isseparated from the second analyte based on charge-to-mass ratio of thefirst analyte relative to the second analyte.